Method and apparatus to provide safety checks for neural stimulation

ABSTRACT

In electrically stimulating neural tissue it is important to prevent over stimulation and unbalanced stimulation which would cause damage to the neural tissue, the electrode, or both. It is critical that neural tissue in not subjected to any direct current or alternating current above a safe threshold. Further, it is important to identify defective electrodes as continued use may result in neural and further electrode damage. Systems and stimulator control mechanisms to prevent damage to neural tissue.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a divisional of application Ser. No. 11/413,771,entitled “Method and Apparatus to Provide Safety Checks for NeuralStimulation”, filed Apr. 28, 2006.

GOVERNMENT RIGHTS NOTICE

This invention was made with government support under grant No.R24EY12893-01, awarded by the National Institutes of Health. Thegovernment has certain rights in the invention.

FIELD OF THE INVENTION

The present invention is generally directed to neural stimulation andmore specifically to an improved method of providing safety checks toprevent neural damage.

BACKGROUND OF THE INVENTION

In 1755 LeRoy passed the discharge of a Leyden jar through the orbit ofa man who was blind from cataract and the patient saw “flames passingrapidly downwards.” Ever since, there has been a fascination withelectrically elicited visual perception. The general concept ofelectrical stimulation of retinal cells to produce these flashes oflight or phosphenes has been known for quite some time. Based on thesegeneral principles, some early attempts at devising a prosthesis foraiding the visually impaired have included attaching electrodes to thehead or eyelids of patients. While some of these early attempts met withsome limited success, these early prosthetic devices were large, bulkyand could not produce adequate simulated vision to truly aid thevisually impaired.

In the early 1930's, Foerster investigated the effect of electricallystimulating the exposed occipital pole of one cerebral hemisphere. Hefound that, when a point at the extreme occipital pole was stimulated,the patient perceived a small spot of light directly in front andmotionless (a phosphene). Subsequently, Brindley and Lewin (1968)thoroughly studied electrical stimulation of the human occipital(visual) cortex. By varying the stimulation parameters, theseinvestigators described in detail the location of the phosphenesproduced relative to the specific region of the occipital cortexstimulated. These experiments demonstrated: (1) the consistent shape andposition of phosphenes; (2) that increased stimulation pulse durationmade phosphenes brighter; and (3) that there was no detectableinteraction between neighboring electrodes which were as close as 2.4 mmapart.

As intraocular surgical techniques have advanced, it has become possibleto apply stimulation on small groups and even on individual retinalcells to generate focused phosphenes through devices implanted withinthe eye itself. This has sparked renewed interest in developing methodsand apparatuses to aid the visually impaired. Specifically, great efforthas been expended in the area of intraocular retinal prosthesis devicesin an effort to restore vision in cases where blindness is caused byphotoreceptor degenerative retinal diseases such as retinitis pigmentosaand age related macular degeneration which affect millions of peopleworldwide.

Neural tissue can be artificially stimulated and activated by prostheticdevices that pass pulses of electrical current through electrodes onsuch a device. The passage of current causes changes in electricalpotentials across visual neuronal membranes, which can initiate visualneuron action potentials, which are the means of information transfer inthe nervous system.

Based on this mechanism, it is possible to input information into thenervous system by coding the information as a sequence of electricalpulses which are relayed to the nervous system via the prostheticdevice. In this way, it is possible to provide artificial sensationsincluding vision.

One typical application of neural tissue stimulation is in therehabilitation of the blind. Some forms of blindness involve selectiveloss of the light sensitive transducers of the retina. Other retinalneurons remain viable, however, and may be activated in the mannerdescribed above by placement of a prosthetic electrode device on theinner (toward the vitreous) retinal surface (epiretial). This placementmust be mechanically stable, minimize the distance between the deviceelectrodes and the visual neurons, and avoid undue compression of thevisual neurons.

In 1986, Bullara (U.S. Pat. No. 4,573,481) patented an electrodeassembly for surgical implantation on a nerve. The matrix was siliconewith embedded iridium electrodes. The assembly fit around a nerve tostimulate it.

Dawson and Radtke stimulated cat's retina by direct electricalstimulation of the retinal ganglion cell layer. These experimentersplaced nine and then fourteen electrodes upon the inner retinal layer(i.e., primarily the ganglion cell layer) of two cats. Their experimentssuggested that electrical stimulation of the retina with 30 to 100 uAcurrent resulted in visual cortical responses. These experiments werecarried out with needle-shaped electrodes that penetrated the surface ofthe retina (see also U.S. Pat. No. 4,628,933 to Michelson).

The Michelson '933 apparatus includes an array of photosensitive deviceson its surface that are connected to a plurality of electrodespositioned on the opposite surface of the device to stimulate theretina. These electrodes are disposed to form an array similar to a “bedof nails” having conductors which impinge directly on the retina tostimulate the retinal cells. U.S. Pat. No. 4,837,049 to Byers describesspike electrodes for neural stimulation. Each spike electrode piercesneural tissue for better electrical contact. U.S. Pat. No. 5,215,088 toNorman describes an array of spike electrodes for cortical stimulation.Each spike pierces cortical tissue for better electrical contact.

The art of implanting an intraocular prosthetic device to electricallystimulate the retina was advanced with the introduction of retinal tacksin retinal surgery. De Juan, et al. at Duke University Eye Centerinserted retinal tacks into retinas in an effort to reattach retinasthat had detached from the underlying choroid, which is the source ofblood supply for the outer retina and thus the photoreceptors. See,e.g., E. de Juan, et al., 99 Am. J. Ophthalmol. 272 (1985). Theseretinal tacks have proved to be biocompatible and remain embedded in theretina, and choroid/sclera, effectively pinning the retina against thechoroid and the posterior aspects of the globe. Retinal tacks are oneway to attach a retinal array to the retina. U.S. Pat. No. 5,109,844 tode Juan describes a flat electrode array placed against the retina forvisual stimulation. U.S. Pat. No. 5,935,155 to Humayun describes aretinal prosthesis for use with the flat retinal array described in deJuan.

SUMMARY OF THE INVENTION

In electrically stimulating neural tissue it is important to preventover stimulation and unbalanced stimulation which would cause damage tothe neural tissue, the electrode, or both. It is critical that neuraltissue in not subjected to any direct current or alternating currentabove a safe threshold. Further, it is important to identify defectiveelectrodes as continued use may result in neural and further electrodedamage. The present invention presents system and stimulator controlmechanisms to prevent damage to neural tissue.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective view of the implanted portion of the preferredretinal prosthesis.

FIG. 2 is a schematic of a circuit for detecting excessive directcurrent flow.

FIG. 3 is a timing diagram of detecting excessive direct current flow.

FIG. 4 is a schematic of a circuit for detecting build up of low leveldirect current flow.

FIG. 5 is a timing diagram of detecting build up of low level directcurrent flow.

FIG. 6 is a schematic of a circuit to detect exceeding a maximum chargeper phase.

FIG. 7 a and 7 b are flowcharts depicting a method for identifyingdefective electrodes.

FIG. 8 a and 8 b are schematic representations of the electrode tissueinterface.

FIG. 9 a-9 e are wave forms illustrating the effects of altering theelectrode tissue interface.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

The following description is of the best mode presently contemplated forcarrying out the invention. This description is not to be taken in alimiting sense, but is made merely for the purpose of describing thegeneral principles of the invention. The scope of the invention shouldbe determined with reference to the claims.

FIG. 1 shows a perspective view of the implanted portion of thepreferred retinal prosthesis. While the invention has broadapplicability to neural stimulation, the preferred embodiment is aretinal prosthesis. A flexible circuit 1 includes a flexible circuitelectrode array 10 which is mounted by a retinal tack (not shown) orsimilar means to the epiretinal surface. The flexible circuit electrodearray 10 is electrically coupled by a flexible circuit cable 12, whichpierces the sclera and is electrically coupled to an electronics package14, external to the sclera.

The electronics package 14 is electrically coupled to a secondaryinductive coil 16. Preferably the secondary inductive coil 16 is madefrom wound wire. Alternatively, the secondary inductive coil 16 may bemade from a flexible circuit polymer sandwich with wire traces depositedbetween layers of flexible circuit polymer. The electronics package 14and secondary inductive coil 16 are held together by a molded body 18.The molded body 18 may also include suture tabs 20. The molded body 18narrows to form a strap 22 which surrounds the sclera and holds themolded body 18, the secondary inductive coil 16, and the electronicspackage 14 in place. The molded body 18, suture tabs 20 and strap 22 arepreferably an integrated unit made of silicone elastomer. Siliconeelastomer can be formed in a pre-curved shape to match the curvature ofa typical sclera. However, silicone remains flexible enough toaccommodate implantation and to adapt to variations in the curvature ofan individual sclera. The secondary inductive coil 16 and molded body 18are preferably oval shaped. A strap 22 can better support an oval shapedcoil.

The preferred prosthesis includes an external portion (not shown) whichincludes a camera, video processing circuitry and an external coil forsending power and stimulation data to the implanted portion.

The electronics package 14 includes an integrated circuit forcontrolling stimulation. The integrated circuit includes an excessivedirect current flow (EDCF) detection circuit as shown in FIG. 2. In theimplant or saline environment, a certain amount of continuous DC currentpassing through an electrode array will cause bubbling which may resultin damage to neural tissue. The EDCF circuit detects harmful DC leakagelevels and transmits this information through the inductive coil 16 tothe external electronics. The integrated circuit in electronics package14 includes a plurality of drivers 202, one for each stimulationelectrode 204, in the electrode array 10. When the external electronicsinitiates an EDCF cycle, switches 206 connect each driver 202 to anelectrode driver common line 208. Any DC leakage current from thedrivers 202 will flow through the electrode driver common line 208. In afault condition, a leakage could flow through the driver output toeither of the power rails, therefore, both sides are tested in alternatestimulation frames. In a particular test, one of the two currents ofthreshold values is turned on to charge the node capacitor 210 of thecommon line. The node capacitor 210 is shown with a dotted line becausethere is no physical capacitor. The node capacitor 210 is inherentcapacitance in the integrated circuit. If there is a leakage path largerthan the threshold value in the opposite direction, the potential of thecommon line is held toward the leakage side. A comparator 212 detectsthe potential against a predetermined middle potential. The output ofthe comparator 212 is captured after a predetermined interval,preferably 250 μs, which is then interpreted by EDCF control logic. Theintegrated circuit further includes power switches such that, in casethere is an EDCF error (i.e. the DC leakage level exceeds the setthreshold, preferably 100 μA), the external electronics will cut offpower to the electrode driver circuits and block the leakage source.FIG. 3, shows the typical timing of the ECDF system. A stimulation cycle302 is followed by the ECDF cycle 304, followed by a shorting cycle 306.Shorting will bleed off any trace amounts of DC build up.

Electrode bubbling occurs when the voltage across the double layer ofthe electrode-tissue interface exceeds a threshold voltage window overcertain time duration during the stimulation cycles. This thresholdvoltage window is found to be around ±1.5V for a flexible circuitelectrode array 10.

Based on the simplified electrical model of the electrode shown in FIG.4, we can see that the charge buildup on the electrode is reflected withthe voltage across the double layer capacitor Cp 402. During normalstimulation with balanced biphasic current pulses, the maximum voltageacross the cap occurs at the end of the first current pulse of amplitudel:

$\left. {{V_{C\mspace{11mu}\max} = {{{{IR}_{p}\left( {1 - {\mathbb{e}}^{{- {Tx}}/\tau_{p}}} \right)} + {{V_{C}(0)}{\mathbb{e}}^{{- {Tx}}/\tau_{p}}}} \approx {{V_{C}(0)} + \frac{Q_{x}}{C}}}}{{{{for}\mspace{14mu} t} ⪡ 1},\mspace{14mu}{{{and}\mspace{14mu} R_{p}} ⪢ 1}}} \right)$

Where Vc(0) is the initial voltage caused by the residue charge left onthe capacitor at the end of shorting, τ_(p)=R_(p)C_(p) (404, 402) is thetime constant of the leaky double layer, Tx is the pulse duration, andQx is the total charge of the first stimulation phase flown through theelectrode.

When a DC leakage is present, the charge is built up even when thesystem is not stimulating. The worst case is when the DC leakage is inthe same polarity as the first phase stimulation current. If the leakagecurrent is small compared to the stimulation current, which may be thecase, the worst case voltage build up on the electrode occurs when thestimulating pulse happens just before the shorting starts. In this case,the maximum electrode voltage is at the start of the second currentpulse, which is illustrated in FIG. 5. The maximum voltage is:V _(Cmax) =I _(L) R _(p) +I _(S) R _(P) e ^(−Tx/τ) ^(p) (1−e ^(−Tx/τ)^(p) )+(V _(C)(0)−I _(L) R _(p))e ^(−(T−T) ^(P7) ^(−Tx)/τ) ^(p)

Where I_(L) is the leakage current, I_(S) is the stimulation current,T_(P7) is P7 profile duration that includes a shorting duration ofT_(SH)=1.4 ms and 0.3 ms for the EDCF check, and T is the stimulationcycle.

A possible electrode voltage map is show in FIG. 5, showing astimulation cycle in the presence of DC leakage that is smaller than thestimulation current. The upper graph 502 is the total current flowingthrough the electrode, and the lower graph 504 is the voltage on thedouble layer capacitor.

The measured parameters C_(p) and R_(p) of the electrodes are preferably0.25-0.3 uF and 70-80K; but 0.25 uF and 80K to handle the worst casecondition, which may result in τ_(p)=20 ms. The shorting time is fixedat 1.4 ms in the preferred embodiment. The stimulation pulse duration Txmay also vary, but the smallest duration is confined by the implantcompliance limit and will make the electrode less tolerant to leakage.From a 15KΩ electrode impedance assumption, this will allow a 0.275 msduration at a current of 400 uA (0.35 mC/cm² maximum charge density).The maximum allowed imbalance of 5% should also be put in the equation.Therefore, in worst case condition, we have:τ_(p)=20 ms, Tx=0.275 ms, I _(S)=400 uA, T _(P7)=1.4 ms, T=8.3 msFrom which we get:V _(Cmax)=0.73V _(C)(0)+21.7I _(L)(mA)+0.43(V)  (3)This voltage should be kept below 1.5V. For a margin of safety, we setthe maximum allowed electrode voltage across the double layer to beV_(Cmax)=1.0V.The steady state value of the residue voltage on the capacitor dependson the shorting duration. It can be estimated as:V _(C)(0)=V _(SH)(0)e ^(−T) ^(SH) ^(/τ) ^(S)   (4)where τ_(S) is shorting time constant, and V_(SH)(0) is the voltage atthe end of the cycle, just before shorting is turned on. In our worstcase condition, we have:V _(SH)(0)=V _(Cmax) e ^(−Tx/τ) ^(p) −(I _(S) −I _(L))R _(p)(1−e^(−Tx/τ) ^(p) )=0.72V _(C)(0)+22.5I _(L)−0.06Where τ_(S)=R_(IR)C_(P)=15K×0.25 uF=3.75 ms.Therefore, with T_(SH)=1.4 ms, we have from (4) and (5):0.73V _(C)(0)−22.5I _(L)+0.06=0  (6)From (3) and (6), we can calculate the maximum allowed asI _(Lmax)=12 μA  (7)This maximum allowable leakage current for an individual electrode toensure that it does not causing electrode bubbling is lower than theEDCF detection circuit (described in reference to FIGS. 2 and 3)threshold current which addresses the total leakage current of allchannels. Therefore a test is needed to detect if the leakage level ofan individual channel passes the maximum allowed leakage value I_(Lmax)estimated.

The electrode potentials takes some time to stabilize because the timeneeded to reach the balance between the charge released by shorting andthe build up by charge imbalance or leakage. The above analysis onelectrode voltages only addresses the steady state condition. Because aDC leakage is considered persistent current flow with or without thepresence of stimulation, a simple method is used to measure theelectrode voltage caused by the leakage current in quiescent conditionin which all stimulation is turned off. This way the disturbance on theelectrode voltage caused by the stimulation currents is avoided. On theother hand, the electrode voltage caused by the leakage current I_(L)(if existent) is still a function of the shorting duration. When we usethe same shorting strategy as the normal condition with EDCF turned off,i.e., T_(SH)=T_(P7=)1.7 ms, the steady state voltages on the electrodedouble layer are estimated:V _(SH−) =V _(SH+) e ^(−(T−T) ^(P7) ^()/τ) ^(p) +I _(L) R _(p)(1−e^(−(T−T) ^(P7) ^()/τ) ^(p) )≈0.718V _(SH+)+22.5I _(L)  (8)V _(SH+) =V _(SH−) e ^(−T) ^(SH) ^(/τ) ^(S) ≈0.635V _(SH−)  (9)Which give us:V _(SH−)=0.50V; V _(SH+)=0.32VWhere V_(SH−) and V_(SH+) are the voltages immediately before and afterthe P7 pulse. Either V_(SH−) or V_(SH+) can be used as the individualelectrode leakage threshold; however, V_(SH−) is preferred because ofits higher value for better accuracy of the measurement. The voltagedriver output V_(DO) is the sum of the I-R drop caused by the real partof the electrode-tissue impedance R_(S) 406 and the electrode voltage onthe double layer capacitor V_(C) discussed above. The I-R drop can becalculated as I_(L)R_(ir), where R_(ir) is the electrode impedance.

From (7), the I-R drop caused by the allowed maximum DC leakage currentis I_(Lmax)R≈0.12-0.56V constant for electrode impedances ranging10-40KΩ. However, when the electrode is lifted from the retina, theimpedance could be lowered to as small as 3KΩ. For a relatively accurateestimation of the leakage current using the V_(DO) measurement, the I-Rdrop should be subtracted from the V_(DO) result using the pre-measuredelectrode impedance values. For simplicity, it may be preferable toignore the I-R drop effects in this leakage detection protocol anddirectly use the V_(DO) data as the electrode voltage V_(C). Ignoringthe I-R drop will yield a 20-50% inaccuracy of the leakage valueestimation that will result in a more conservative monitoring. This willnot compromise the safety but will simplify the measurement.

In our electrode impedance measurement protocol, an electrode withimpedance 65KΩ or higher is labeled as an open electrode. Stimulation toan open electrode is turned off. However, it must be noted that anelectrode with impedance higher than 65KΩ could still bubble because ofDC leakage. Assume that the maximum allowed leakage is still 12 μA andthe compliance limit for the leakage current is 7.0V, we can calculatewith the same method used above that an electrode with impedance valueas high as 450KΩ will reach electrode bubbling status. However, themeasured voltage is limited to about 4V on the anodic side, which willlimit the detectable threshold leakage to electrodes with impedance upto 230KΩ. For example, if the measured electrode impedance is 200KΩ,then a measured V_(DO) of 3.4V or more can be considered a bubblingstatus. In either case, the electrode impedance measurement should beable to discriminate between a high impedance electrode and an openelectrode.

The following method is suggested to label an electrode as a highimpedance (HI) electrode: After the regular impedance measurementroutine, all “open” electrodes (if any) are measured again using 8.1μA/1 ms current pulses with the V_(DO) sampling point set at 0.9 msafter the pulse start. An electrode with measured impedance 500KΩ orless (V_(DO)<±4V) shall be tagged as HI electrode, while higherimpedance electrodes are tagged as open.

For HI impedance electrodes, the build up voltage on the electrodecapacitor is solely from DC leakage because they are not stimulated.Therefore, the maximum allowed electrode voltage is simply V_(SH−)=1V,plus the I-R component when using V_(DO) measurement. Therefore, thethreshold for an HI electrode should be V_(DO) (measured at V_(SH−)):VDO| _(V) _(SH−) <=I _(Lmax) R+1≈0.012R(KΩ)+1  (10)

Considering the limited ADC range in the anodic direction, it may bepreferable to check all HI electrodes with impedance 200KΩ or higheragainst the threshold voltage of 200KΩ, which is 3.4V. For openelectrodes, this checking procedure can be omitted.

Referring to FIG. 6, the problem of exceeding maximum charge per phase(MCPP) can be solved with two different methods. In one case, thecathodic and anodic pulse amplitude of each electrode is multiplied bythe phase duration and the result is checked against a predeterminedMCPP value. If the value is exceeded, the amplitude value is clipped toachieve the limit. This first method is accurate for square waves, butbecomes less accurate as the stimulation wave form becomes arbitrary. Inanother method, a discrete integral is taken in the time representationof the pulse to calculate the charge. If the integral at any point oftime exceeds the absolute MCPP value (positive or negative), theamplitude value is either clipped to achieve the limit or zeroed out.Another method that could be used is by an active current monitor on thestimulator device, which limits the current to a particular MCPP. Forall these methods, the implementation will be stimulator settingdependant (e.g. Frequency setting) so that the safety checks will adaptto limit the charge to the MCPP in any case.

The circuit shown in FIG. 6, performs an integration of the waveformthrough capacitance charge. The stimulation signal 610 is amplified intransistors 612 to drive the stimulating electrode 614. Simultaneously,the stimulation signal 610 is amplified in transistors 616 to create amuch smaller, but proportional charge signal. Capacitor 618, between thecharge signal output and the common electrode 620 is charged by thecharge signal. Comparator 622 compares the charge on capacitor 618 witha predetermined maximum charge 624 and generates an over limit signalwhen the charge on capacitor 618 is over the predetermined maximumcharge.

Another way to limit current density is to provide a compromise betweenthe monopolar mode and bipolar mode of stimulation. The method includessetting up stimulation wave forms as in bipolar stimulation mode withoutdisconnecting the common electrode (ground). A portion of the currentwill flow between the electrodes, and a portion of the current will flowbetween the electrodes and ground. The ratio of current flow will dependon electrode impedance. This hybrid bipolar mode of stimulation couldpossibly result in lower thresholds than can be obtained from a truebipolar mode of stimulation, and also has the advantage of greaterselectivity than can be obtained from a monopolar mode of stimulation.To set up a safe hybrid bipolar stimulation wave form, the twoelectrodes concerned should have balanced biphasic currents going inopposite phases at exactly the same times, with the common electrodeconnected. If the pulse waveforms are balanced by themselves on eachelectrode but overlap non-contiguously in opposite phases with otherelectrodes in the array, a resultant unbalanced current could flowthrough the tissue eventually causing neural tissue damage. A safetycheck method could be implemented in the external electronics to preventsuch unbalanced multipolar waveforms to be sent to the stimulator. Theabove phenomenon and safety check method is also applicable tomultipolar forms of stimulation (i.e. in addition to bipolar).

It is also important to limit the maximum stimulation across allelectrodes in an electrode array. While each electrode is individuallystimulating at a safe level, there can still be neural damage or even insome cases pain if all electrodes are stimulating at or near theirindividual maximum level. Hence, it is important to track and limit thesum total stimulation from all electrodes. This is a simple calculationthat can be done with software in the external electronics. For eachstimulation cycle, all stimulation values are summed and compared with apredetermined maximum. If the sum exceeds the predetermined maximum, thestimulation is reduced, either proportionally across all electrodes, orby limiting electrodes set to higher stimulation levels.

Broken electrode detection can be achieved with a method of monitoringimpedance of the electrodes (over time (delta l), as well as comparingimpedances of electrodes to its neighboring electrodes in the media, andincorporation of physiological data as observed from physicians). When abroken electrode is found in the system, the stimulator device iscommanded to halt stimulation on that broken electrode—with the historyof electrode damage logged and persistent in the stimulator controller.Visually, the electrode health is represented through acolor/topographical map of the electrode array on a computer system.Optimally, a movie-like data playback of the impedances can be provided.Electrode damage can also consist of shorting between electrodes (asopposed to ‘broken’) which will also be detected through impedancemonitoring.

Charge imbalance can be reduced by implementing a 1.4 ms (adjustable, asmentioned in (3)) shorting pulse prior to any stimulation on everystimulation frame of the stimulator. Additionally, the stimulatorcontroller can adjust (and check) the anodic and cathodic current level,using amplitude parameter tables determined from the manufacturing testsof the stimulator, to achieve the best balance. The stimulatorcontroller also checks that there exist no overlapping anodic/cathodicprofiles that could cause a charge imbalance at the tissue. Anothermethod of reducing charge imbalance on an electrode is by having thestimulator controller auto-balancing the pulse. Two ways of achievingthis are by appending a square pulse after the pulse to compensate forthe imbalance or by appending an inverted pulse to compensate. (5) Checkagainst DC flow in tissue from the electrodes is achieved through acombination of the ASIC test, the shorting function, and an initialindividual electrode check implemented on startup (or periodically) bythe stimulator controller. (6) The issue of implant overheating ishandled in multiple ways. Power is controlled on the retinal prosthesisthrough a feedback loop fed by implant back telemetry. Currently, thisis achieved by inferring the implant heat through the current in theshunt regulators of the implant device. Additionally, if a thermistor isplaced in the implant device, then the heat can be measured and returnedto the controller through the back telemetry link.

Additionally, the instantaneous current output on the stimulator deviceis limited to a constant value by the controller. This ensures that noamount of instantaneous current is allowed that might expect thestimulator to reset due to lack of power. This instantaneous currentlimit could also be variable (instead of constant) with appropriate backtelemetry and controller design. For system operational modes that donot allow heat control via the shunt current values, the controllerperforms a check that ensures the implant shunt current level can be setto a safe value immediately prior to the operational mode which doesn'tprovide shunt current information in the back telemetry. Also, thestimulator controller verifies the voltage setting with an ADC circuitupon any change of voltage to the RF power circuitry.

Referring to FIG. 7 a, the electrode failure detection system includes atable of electrode impedance value that is N, the number of electrodes,by T, the time history stored, and a table of electrode state valueswith three states, good, questionable, and bad. Each test begins byloading a time value 702 to identify the test. The electrode counter isset to zero 704, and the first electrode is tested 706. The systemdetermines if the electrode impedance is within hard limits 708. Theselimits are the same for all electrodes. This can be used to catch themost obvious failures. As an illustration, the impedance value returnedcan range from 0Ω (complete short) to 65,535Ω (complete open), where thenominal value for a good electrode might be in the range of 10,000Ω to20,000Ω. Here, a lower limit of 2,000Ω and an upper limit of 50,000Ωcould be used, where any electrode falling outside these ranges woulddefinitively be experiencing a failure and will be marked bad 710. Ifthe electrode is within the hard limits, the electrode impedance isstored in the table 712. The process is repeated 714 incrementing theelectrode counter 716 each time.

Once all electrodes have been tested, the system calculates and stores amedian and standard deviation 718. If this is the test has beenpreviously preformed 720, the impedance data is shifted to the nextmemory location 722. If not, calculate the standard deviation of goodelectrodes 724, and shift the standard deviation to the previous memorylocation 726. If the electrode values are close, a very low standarddeviation may trigger too many electrodes marked as bad. Hence thestandard deviation is compared with a preset minimum 728 and replacedwith the minimum if the minimum is higher 722. If the electrode valuesare far apart, a very high standard deviation may not mark badelectrodes. Hence the standard deviation is compared with a presetmaximum 730 and replaced with the maximum if the maximum is lower 732.

It has been observed that an electrode which fails (reading as high as65 kΩ initially) may give subsequent impedance values which begin tofall, sometimes returning to a nominal, good value. This is apparentlydue to fluid leaking in the broken area and creating an alternative (andundesirable) conduction path. Thus, any one set of readings may not showall electrodes known obviously to be broken, and a history must bemaintained. If an electrode ever exceeds the limits, it is consideredpermanently failed.

Note the history of the electrodes should not begin until theimplantation of the array has stabilized. Impedance readings are veryuseful through the implantation process, but values may shiftdrastically, and false failures can impede the process.

It appears many failed electrodes never reach a high value, or reach itand return so quickly it may not be recorded. Thus, additional checksare required.

Referring to FIG. 7 b, the electrode counter is again set to zero 740.If the electrode is already marked bad 742, it cannot be marked goodagain. For each electrode if the absolute value of the electrodeimpedance minus the median is greater than an unsafe limit times thestandard deviation 744, the electrode marked bad 746. In the preferredembodiment, the unsafe limit is four times the standard deviation. Ifthe absolute value of the electrode impedance minus the median isgreater than a questionable limit times the standard deviation 748 andthe electrode has measured questionable a predetermined number of timepreviously 750, the electrode is marked as bad 746. In the preferredembodiment, the questionable limit is three times the standarddeviation. The process is repeated until all electrodes have beenchecked 754, with the electrode counter incremented each time 756.Finally, the standard deviation is recalculated for the remaining goodelectrodes 758 and stored 760.

In the preferred embodiment, impedance is used to determine theelectrode integrity at the interface of the neural stimulator andtissue, as well as the integrity of the electrode stimulation paththrough the implantable device. However, data obtained both in vitro andin vivo show that these impedance readings are time dependent, and anygiven temporal snapshot may show failed electrodes as having perfectlynominal values. Furthermore, potentially critical failures such aselectrode movement or loss of tissue contact may not be quicklydiagnosed using the current methods as the changes in impedance may notbe large.

Capacitance measurement is a possible solution since electro-neuralinterfaces have distinct capacitive characteristics. Referring to FIG. 8a an interface between an electrode and tissue is actually twointerfaces, one from electrode to fluid 802 and a second from fluid totissue 804 as show in FIG. 8. However, when an electrode lifts offtissue, the resulting circuit resembles a single electrode fluidinterface 806 as shown in FIG. 8 b. When an electrode is making goodtissue contact, it follows the tissue model shown in FIG. 8 a, with anaverage value of 15 k-25 kΩ. When an electrode has completelylifted-off, it follows the fluid model shown in FIG. 8 b. It isdesignated ‘fluid’ since the values & characteristics of an implantedarray lifted up in fluid are the same as that in saline in thelaboratory.

When an electrode begins to lift off, it begins making contact with thefluid, and thus has two parallel impedance paths. A simple impedancemeasurement will show only a slight drop at the beginning. However, thephase characteristics of the two paths are different, both in totalcapacitance value and total phase shift. Thus, the waveform should showthe effect.

Referring to FIG. 9 for a simple model on how a capacitor would affectan AC square wave. Initially it may appear simple to measure V₂ and V₁and establish ΔV as V₂−V₁ and thus a measure of capacitance, as ΔV wouldincrease as the total capacitance in the circuit increases. However, asgiven in the schematic of a typical electrical model of theelectrode-tissue interface, FIG. 8, the portion of electricity goingthrough the capacitor is phase shifted by 90°. The resultant waveformdoes show a capacitive phase shift, but with rounded peaks potentiallyanywhere on the waveform, so ΔV would not be a measure of capacitance. Amethod is required that accounts for this phase shift which couldreflect the total capacitance as well as changes in its characteristics.

The difference in capacitance between the various electrical paths whichoccurs in both healthy and failed electrodes will be manifested as aphase difference in the various waves that sums the resultant measuredwaveform. This difference can be computed for each stimulation phase,based on the area of the curve of the first derivative, subtracting thebaseline waveform where no phase shift is present. Healthy electrodesappear to have distinctive measurable characteristics in both anodic andcathodic phases. These characteristics change significantly when anelectrode is losing contact at the neural interface (array lift-off) oris degrading, and thus electrodes which show a significant difference inphase shift are likely experiencing the manifestation of a failure. Thismethod could be used to detect array lift-off potentially before itwould be detected by the associated impedance values currently in use.

Accordingly, what has been shown is an improved method of stimulatingneural tissue for improved response to brightness. While the inventionhas been described by means of specific embodiments and applicationsthereof, it is understood that numerous modifications and variationscould be made thereto by those skilled in the art without departing fromthe spirit and scope of the invention. It is therefore to be understoodthat within the scope of the claims, the invention may be practicedotherwise than as specifically described herein.

The invention claimed is:
 1. A method of stimulating retinal neuraltissue comprising: providing a neural stimulator including a pluralityof electrodes suitable to stimulate retinal neural tissue; providing abiphasic square wave pulse to each electrode; calculating capacitiveimpedance by a first derivative of the voltage across a stimulationpulse; testing the impedance of all electrodes by obtaining voltagemeasurements from all electrodes in response to said square wave pulses;recording changes of capacitive impedance over time derived from saidvoltage measurements; and determining lifting of electrodes based onchanges of a capacitive component of said impedance by considering phaseshifts in said changes of impedance over time; wherein the changes ofthe capacitive component are computed based on a phase differencederived from a difference between a resultant voltage waveform obtainedfrom the voltage measurements and a baseline voltage waveform withoutphase shifts derived from said voltage measurements.
 2. The methodaccording to claim 1, further comprising discontinuing stimulation onsaid lifted electrodes.
 3. The method of claim 1, wherein the changes ofthe capacitive impedance are computed in part based on a phasedifference derived from a difference between an area under a curve ofthe resultant voltage waveform and the baseline voltage waveform withoutphase shifts.
 4. The method of claim 1, wherein the changes of thecapacitive impedance are computed in part based on a phase differencederived from a difference between an area under a curve of a firstderivative of the resultant voltage waveform and the baseline voltagewaveform without phase shifts.
 5. The method of claim 1, wherein thebaseline voltage waveform is subtracted from the resultant voltagewaveform.
 6. The method of claim 1, wherein determining liftingelectrode further includes comparing the phase shift of anodic andcathodic phases of the baseline voltage waveform and the resultantvoltage waveform.